Field of the Invention
The present invention relates to bioresorbable scaffolds; more particularly, this invention relates to bioresorbable scaffolds for treating peripheral vessels of the body.
Description of the State of the Art
Radially expandable endoprostheses are artificial devices adapted to be implanted in an anatomical lumen. An “anatomical lumen” refers to a cavity, duct, of a tubular organ such as a blood vessel, urinary tract, and bile duct. Stents are examples of endoprostheses that are generally cylindrical in shape and function to hold open and sometimes expand a segment of an anatomical lumen (one example of a stent is found in U.S. Pat. No. 6,066,167 to Lau et al). Stents are often used in the treatment of atherosclerotic stenosis in blood vessels. “Stenosis” refers to a narrowing or constriction of the diameter of a bodily passage or orifice. In such treatments, stents reinforce the walls of the blood vessel and prevent restenosis following angioplasty in the vascular system. “Restenosis” refers to the reoccurrence of stenosis in a blood vessel or heart valve after it has been treated (as by balloon angioplasty, stenting, or valvuloplasty) with apparent success.
The treatment of a diseased site or lesion with a stent involves both delivery and deployment of the stent. “Delivery” refers to introducing and transporting the stent through an anatomical lumen to a desired treatment site, such as a lesion. “Deployment” corresponds to expansion of the stent within the lumen at the treatment region. Delivery and deployment of a stent are accomplished by positioning the stent about one end of a catheter, inserting the end of the catheter through the skin into an anatomical lumen, advancing the catheter in the anatomical lumen to a desired treatment location, expanding the stent at the treatment location, and removing the catheter from the lumen.
In the case of a balloon expandable stent, the stent is mounted about a balloon disposed on the catheter. Mounting the stent typically involves compressing or crimping the stent onto the balloon prior to insertion in an anatomical lumen. Crimping refers to an iris-type or other form of crimper, such as the crimping machine disclosed and illustrated in US 2012/0042501. At the treatment site within the lumen, the stent is expanded by inflating the balloon. The balloon may then be deflated and the catheter withdrawn from the stent and the lumen, leaving the stent at the treatment site. In the case of a self-expanding stent, the stent may be secured to the catheter via a retractable sheath. When the stent is at the treatment site, the sheath may be withdrawn which allows the stent to self-expand.
The stent must be able to satisfy a number of basic, functional requirements. The stent must be capable of withstanding the structural loads, for example, radial compressive forces, imposed on the stent as it supports the walls of a vessel after deployment. Therefore, a stent must possess adequate radial strength. After deployment, the stent must adequately maintain its size and shape throughout its service life despite the various forces that may come to bear on it. In particular, the stent must adequately maintain a vessel at a prescribed diameter for a desired treatment time despite these forces. The treatment time may correspond to the time required for the vessel walls to remodel, after which the stent is no longer necessary for the vessel to maintain a desired diameter.
A commonly used type of peripheral stent is the self-expanding stent made from super-elastic material, such as Nitinol. This type of material is known for its ability to return to its original configuration after severe deformation, such as a crushing load or longitudinal bending. However, this variety of self-expanding stents have undesired qualities; most notably, the high resiliency of super-elastic material produces what is commonly referred to as a “chronic outward force” (COF) on the blood vessel supported by the stent. Complications resulting from COF are discussed in Schwartz, Lewis B. et al. Does Stent Placement have a learning curve: what mistakes do we as operators have to make and how can they be avoided?, Abbott Laboratories; Abbott Park, Ill., USA. It is believed that a COF exerted on a blood vessel by a self-expending stent is a main contributor to high degrees of restenosis of lesions treated by the self-expanding stent. It has been shown that not even an anti-proliferative drug delivered from drug eluting self-expandable stents can mitigate the restenosis caused by the stent's COF. Stents that are plastically deformed by a balloon to support a vessel do not suffer from this drawback. Indeed, balloon expanded stents, in contrast to self-expanding stents made from a super-elastic material, have the desirable quality of being deployable to the desired diameter for supporting the vessel without exerting residual outward forces on the vessel.
A balloon-expanded polymer scaffold, such as that described in US 2010/0004735 is made from a biodegradable, bioabsorbable, bioresorbable, or bioerodable polymer. The terms biodegradable, bioabsorbable, bioresorbable, biosoluble or bioerodable refer to the property of a material or stent to degrade, absorb, resorb, or erode away from an implant site. The polymer scaffold described in US 2010/0004735, for example, as opposed to a metal stent, is intended to remain in the body for only a limited period of time. In many treatment applications, the presence of a stent in a body may be necessary for a limited period of time until its intended function of, for example, maintaining vascular patency and/or drug delivery is accomplished. Moreover, it has been shown that biodegradable scaffolds allow for improved healing of the anatomical lumen as compared to metal stents, which may lead to a reduced incidence of late stage thrombosis. In these cases, there is a desire to treat a vessel using a polymer scaffold, in particular a bioabsorable or bioresorbable polymer scaffold, as opposed to a metal stent, so that the prosthesis's presence in the vessel is for a limited duration. However, there are numerous challenges to overcome when developing a polymeric scaffold.
The art recognizes a variety of factors that affect a polymeric scaffold's ability to retain its structural integrity and/or shape when subjected to external loadings, such as crimping and balloon expansion forces. These interactions are complex and the mechanisms of action not fully understood. According to the art, characteristics differentiating a polymeric, bioresorbable scaffold of the type expanded to a deployed state by plastic deformation from a similarly functioning metal scaffold are many and significant. Indeed, several of the accepted analytic or empirical methods/models used to predict the behavior of metallic scaffolds tend to be unreliable, if not inappropriate, as methods/models for reliably and consistently predicting the highly non-linear, time dependent behavior of a polymeric load-bearing structure of a balloon-expandable scaffold. The models are not generally capable of providing an acceptable degree of certainty required for purposes of implanting the scaffold within a body, or predicting/anticipating the empirical data.
Polymer material considered for use as a polymeric scaffold, e.g. poly(L-lactide) (“PLLA”), poly(L-lactide-co-glycolide) (“PLGA”), poly(D-lactide-co-glycolide) or poly(L-lactide-co-D-lactide) (“PLLA-co-PDLA”) with less than 10% D-lactide, and PLLD/PDLA stereo complex, may be described, through comparison with a metallic material used to form a stent, in some of the following ways. A suitable polymer has a low strength to volume ratio, which means more material is needed to provide an equivalent mechanical property to that of a metal. Therefore, struts must be made thicker and wider to have the required strength for a stent to support lumen walls at a desired radius. The scaffold made from such polymers also tends to be brittle or have limited fracture toughness. The anisotropic and rate-dependent inelastic properties (i.e., strength/stiffness of the material varies depending upon the rate at which the material is deformed, in addition to the temperature, degree of hydration, thermal history) inherent in the material, only compound this complexity in working with a polymer, particularly, bioresorbable polymer such as PLLA or PLGA.
Processing steps performed on, and design changes made to a metal stent that have not typically raised concerns for, or required careful attention to unanticipated changes in the average mechanical properties of the material, therefore, may not apply to a polymer scaffold due to the non-linear and sometimes unpredictable nature of the mechanical properties of the polymer under a similar loading condition. It is sometimes the case that one needs to undertake extensive validation before it even becomes possible to predict more generally whether a particular condition is due to one factor or another—e.g., was a defect the result of one or more steps of a fabrication process, or one or more steps in a process that takes place after scaffold fabrication, e.g., crimping. As a consequence, a change to a fabrication process, post-fabrication process or even relatively minor changes to a scaffold pattern design must, generally speaking, be investigated more thoroughly than if a metallic material were used instead of the polymer. It follows, therefore, that when choosing among different polymeric scaffold designs for improvement thereof, there are far less inferences, theories, or systematic methods of discovery available, as a tool for steering one clear of unproductive paths, and towards more productive paths for improvement, than when making changes in a metal stent.
The present inventors recognize, therefore, that, whereas inferences previously accepted in the art for stent validation or feasibility when an isotropic and ductile metallic material was used, those inferences would be inappropriate for a polymeric scaffold. A change in a polymeric scaffold pattern may affect not only the stiffness or lumen coverage of the scaffold in its deployed state supporting a lumen, but also the propensity for fractures to develop when the scaffold is crimped or being deployed. This means that, in comparison to a metallic stent, there is generally no assumption that can be made as to whether a changed scaffold pattern may not produce an adverse outcome, or require a significant change in a processing step (e.g., tube forming, laser cutting, crimping, etc.). Simply put, the highly favorable, inherent properties of a metal (generally invariant stress/strain properties with respect to the rate of deformation or the direction of loading, little effect of body temperature or hydration, and the material's ductile nature), which simplify the stent fabrication process, allow for inferences to be more easily drawn between a changed stent pattern and/or a processing step and the ability for the stent to be reliably manufactured with the new pattern and without defects when implanted within a living being.
A change in the pattern of the struts and rings of a polymeric scaffold that is plastically deformed, both when crimped to, and when later deployed by a balloon, unfortunately, is not predictable to the same or similar degree as for a metal stent. Indeed, it is recognized that unexpected problems may arise in polymer scaffold fabrication steps as a result of a changed pattern that would not have necessitated any changes if the pattern was instead formed from a metal tube. In contrast to changes in a metallic stent pattern, a change in polymer scaffold pattern may necessitate other modifications in fabrication steps or post-fabrication processing, such as crimping and sterilization.
Scaffolds used to treat coronary vessels experience, for the most part, a primarily radial loading. However, scaffolds intended for peripheral vessels experience a quite different loading, to such an extent that the traditional measure of a stent's fitness for use, i.e., its radial strength/stiffness, is not an accurate measure of whether the scaffold will have sufficient strength to provide mechanical support within the peripheral vessel for the duration needed. This is because a peripheral scaffold is placed in a significantly different environment from a coronary scaffold. The vessel size is larger. And there is much more movement of the vessel, with motions in different directions, especially when located close to an articulating joint. As such, a scaffold intended for a peripheral vessel will need to be able to sustain more complex loading, including a combination of axial, bending, torsional and radial loading. See e.g. Bosiers, M. and Schwartz, L., Development of Bioresorbable Scaffolds for the Superficial Femoral Artery, SFA: CONTEMPORARY ENDOVASCULAR MANAGEMENT (‘Interventions in the SFA” section). These and related challenges facing peripherally implanted stents and scaffolds are also discussed in U.S. application Ser. No. 13/015,474.
The art has previously addressed a need to reduce loading on ring elements due in-part to axial, bending, torsion or combination of these loads when the scaffold is implanted within a peripheral vessel. Previously proposed designs introduce frangible or significantly reduced stiffness in connecting links between ring elements, as in for example US 2011/0190972, or by forming a scaffold composed of several disconnected scaffold segments, as in U.S. application Ser. No. 13/840,257.
A continued need exists for addressing the loading environment of a peripherally-implanted scaffold, such as a scaffold implanted in the superficial femoral-artery including, but not limited to axial or longitudinally loading which can lead to fracture or weakening of struts forming ring elements; providing a coupling between scaffold segments while, at the same time, avoiding the creation of load paths into ring structures; and/or providing a coupling for maintaining a desired spacing between scaffold segments without losing any of the benefits of having a de-coupled or separate scaffold segments that in the aggregate form the scaffold for treating a diseased peripheral vessel.